The listing or discussion of an apparently prior-published document in this specification should not necessarily be taken as an acknowledgement that the document is part of the state of the art or is common general knowledge.
There are many medical situations where it is necessary or desirable to implant a stent within a patient in order to prevent or counteract a constriction in a naturally occurring vessel or passage. In this context, a “stent” is an artificial tubular structure which is able to apply force radially outwardly on a vessel or passage of a patient in order to maintain patency of the vessel or passage and permit fluid flow through said vessel or passage.
One of the main uses of stents is in the treatment of cardiovascular disease, which is a leading cause of mortality within the developed world. Coronary disease is of most concern and patients having such disease usually have narrowing in one or more coronary arteries. One treatment is coronary stenting, which involves the placement of a stent at the site of acute artery closure. This type of procedure has proved effective in restoring vessel patency and decreasing myocardial ischemia.
Stents are also commonly used in the treatment of other conditions caused by the narrowing of the vasculature, for example, peripheral arterial disease and renal vascular hypertension.
Current stent technology is based on the use of permanent stents made from corrosion-resistant metals, such as 316L stainless steel, or metal alloys, such as cobalt chromium or nitinol. The inherent strength of metals mean that stents made from such metallic tubes can adopt a low profile while exhibiting the radial strength needed to maintain vessel patency (i.e. keeping the vessel in an open and unobstructed state) whilst retaining a low profile. The profile of a stent is to be understood as relating to its physical dimensions, in particular, its wall thickness and diameter.
However, despite their low profile and radial strength characteristics, there are a number of disadvantages associated with the use of permanent metallic implants. In particular, exposure of the currently used metallic stents to flowing blood can result in thrombus formation, smooth muscle cell proliferation and acute thrombotic occlusion of the stent. Furthermore, metallic stents have specific drawbacks which limit their widespread use throughout the body. These limitations include long-term endothelial dysfunction, delayed re-endothelialisation, thrombogenicity, permanent physical irritation, chronic inflammatory local reactions, mismatches in mechanical behaviour between stented and non-stented vessel areas, inability to adapt to growth in young patients, and importantly non-permissive or disadvantageous characteristics for later surgical revascularization.
The major effect of stent implantation is provided by its scaffolding effect, which is required to last for between 6 to 12 months, during which time vessel patency can be restored to near normal levels. After this period of time, the presence of a stent within the vessel does not usually provide any beneficial effects in the long term as regards its role as a supporting structure.
In the light of the disadvantages associated with permanent metallic stents, the general consensus amongst medical practitioners over recent years has been the desire to move away from using permanent stents and towards using non-permanent biodegradable stents.
In order for the use of biodegradable stents to be realised in a clinical setting, they must possess the following: (1) Mechanical strength—the biodegradable stent must exhibit mechanical strength approaching that of metallic stents so that it can retain a low profile while at the same time being able to withstand the radial pressures exerted upon it in the vessel environment; (2) Optimum degradation profile—the stent must remain in place and maintain its structural integrity long enough for vessel patency to be restored. However, once the task of supporting the vessel has been achieved, degradation of the stent needs to be reasonably swift so as to prevent the onset of any unwanted side-effects. It should be noted that this balancing act is not as trivial as it first appears; and (3) Biocompatibility—the degradation products of many bioabsorbable compounds are capable of eliciting inflammatory immune responses. Therefore, the materials comprising the stent and their degradation products must be biocompatible in that they do not elicit such responses.
Many biodegradable stents are undergoing development and a number of fully biodegradable stents are currently being examined in a number of clinical trials. In addition to adopting purely a support role within the vessel, many biodegradable stents are also designed to be drug eluting. Such stents have been assessed in clinical trials and include Abbot's BVS Stent (Ormiston J. A., et al. Lancet, 2008, 371, p 899-907) and Biotronic's Magnesium Stent (Erbel R., et al. Lancet, 2007, 369, p 1869-1875). By way of a specific example, Abbott's BVS stent is fabricated from a biodegradable polyester derived from lactic acid (poly-L-lactic acid, PLLA) with a coating that controls release of the drug everolimus to prevent rejection and reclogging. An example of a non drug eluting biodegradable stent is Igaki Medical's Igaki-Tamai's stent (Tamai H., et al. Circulation, 2000, 102, p 399-404), which is also fabricated from PLLA.
In order for stents to function effectively they must have a radial strength capable of withstanding the radial compressive forces exerted by the luminal wall of a blood vessel. Moreover, they must exhibit sufficient flexibility to allow for crimping onto a balloon catheter for the journey through the tortuous vascular network to the site of deployment and for expansion at said site of deployment.
Temporary stents have been made from biodegradable metallic tubing, for example, Biotronic's Magnesium Stent. This stent is a tubular, slotted stent sculpted by laser from a tube of a biodegradable magnesium alloy. Like the permanent stainless steel stents, it has low elastic recoil, with minimum shortening after inflation. Despite having these properties, only limited success has been observed in clinical trials. This has partly been attributed to the relatively rapid rate of degradation of 60 to 90 days.
Given the problems associated with the rapid degradation of such biodegradable metallic stents, stents made from biodegradable polymeric tubing are attractive. Firstly, a myriad of polymeric materials are already known in the art as compared to biodegradable metals, which are essentially limited to the use of magnesium and iron.
Secondly, the degradation rates of biodegradable polymers are in the range of months and years and so are generally slower than that of biodegradable metals, which are generally measured in weeks. Furthermore, it is possible to alter the degradation rate of a polymeric material to suit specific needs by adjusting the composition of the polymer or polymer blend used. However, despite these advantages, there are a number of problems that need to be overcome in order to make stents fashioned from biodegradable polymers a viable alternative to metallic stents.
The inherent properties of metals mean that they are ideal for producing low profile stents exhibiting the radial strength required to maintain the lumen of the blood vessel open. Compared to metals and metal alloys, polymers have an inferior strength to weight ratio. Therefore, if a polymeric stent is compared to a metallic stent having a similar slot/mesh size and strut/wall thickness it would be lacking in the mechanical strength required to withstand the radial forces exerted upon it by a blood vessel wall. There are various solutions to compensate for this strength differential, however none are ideal.
The radial strength of a polymeric stent can be increased by reducing the cell size of the mesh. However, the problem with decreasing the cell size is that the flexibility of the stent is reduced, which can make implantation of the stent difficult because blood vessels are not perfectly cylindrical in shape and thus the natural conformation of a blood vessel may be lost when the stent is implanted.
The radial strength can also be improved by increasing the thickness of the stent wall struts. However, this increases the profile of the stent and there is evidence that suggests that having thicker struts in a mesh stent results in a greater likelihood of restenosis after implantation of the stent.
Therefore, it would be highly desirable to produce tubing made from a biodegradable material with mechanical strength characteristics such that it could be fashioned into a stent having a similar strut and mesh size common amongst permanent metallic stents that are currently used in the clinic.
Polymeric tubing formed by extruding a polymer melt from, for example, a single or twin screw extruder, exhibits minimal alignment of the polymer molecules. Alignment of these molecules in both the radial and axial directions improves the overall properties of the tubing. A number of techniques, such as blow molding and die drawing, can deform polymeric tubing so as to induce molecular orientation of the polymer molecules, in either a uni- or biaxial fashion, thus strengthening said tubing.
In the context of polymer tubing, blow molding is a process whereby a tube, fixed at both ends by some form of grip and held within a cylindrical mould, is heated to a temperature between its glass transition and melting temperature. To achieve the target diameter gas is then pumped through the heated tubing to push the walls of the tubing against the boundary created by the mold.
Blow molding has previously been used to manufacture polymeric tubing for use in biodegradable stents. For examples of blow molding techniques, see US 2010/00258894 A1, US 2010/0198331 A1, U.S. Pat. No. 7,971,333 B2 and US 2011/0062638 A1. Given the nature of these blow molding techniques, they are unable to produce tubing with the size required for stents in a continuous manner. Furthermore, there is a considerable amount of waste material retained in the fixing means.
Die drawing is a process whereby a polymeric material is heated to a temperature between its glass transition and melting temperature, and pulled through a die to change its cross sectional area. The deformation during this change in cross-sectional area causes orientation and alignment of the polymer molecules which gives improvements in terms of strength and stiffness. Unlike blow molding techniques, die drawing can also produce tubing in a continuous manner because the process does not require the tubing to be fixed at both ends. However, die drawing has never been used to produce tubes capable of being used for stents.
In U.S. Pat. No. 4,801,419 a die drawing process was used to produce oriented polymeric tubing. In one example, a length of unplasticised PVC thick walled tubing having an inner diameter of 32 mm and an outer diameter of 42 mm was drawn over the expanding cone of a mandrel and through a die, to give a die drawn tube with a wall thickness of 3.7 mm. Similarly, tubing having a wall thickness of 0.225 mm was produced in U.S. Pat. No. 5,650,114, by deforming a tube over an expanding former (mandrel). The resultant tubes produced by these methods are useful in the fields of gas piping etc., but are not suitable for use in the manufacture of stents due to their large size.
Die drawing is a thermal process. Therefore, scaling such processes, in particular, down-scaling, is non-trivial due to the differences in volumes, surface areas and heat transfer rates involved. The polymeric materials that are used in bioresorbable stents are highly temperature and moisture sensitive which adds to the difficulty of producing tubing suitable to use in stents by die drawing. This is in contrast to the conventional pipe grade plastic used in the above mentioned die drawing processes.
Typically, stents are manufactured from polymeric tubes by using a laser to cut away the wall of the tube to create the required mesh-like scaffolding structure of a stent. As laser cutting can be particularly sensitive to fluctuations in the thickness of the tube wall the tube must have a uniform shape and consistent wall thickness along its length for the process to be successful. While blow molding can achieve the required uniform dimensions without significant difficulties, uniformity is difficult to attain with die drawing techniques.
In view of the above, a die drawing process that consistently produces tubing having the dimensions suitable for use in a stent, i.e. a wall thickness of less than 150 microns and an outer diameter of 1-3 mm, would be useful as no such technique has been disclosed.
Therefore, to address the above-mentioned problems the present inventors have devised a die drawing method for the production of polymeric tubing for use in stents, said tubing having optimal, or otherwise improved, mechanical strength and shape characteristics.
The above discussion has focussed on tubing for use in coronary, peripheral, cardiothoracic, and neuro vascular stents but it is to be understood that the present invention is not limited thereto. Tubing for stents other than vascular stents, such as tubing for ureteral, urethral, duodenal, colonic and biliary stents are also relevant to the current invention.